In MRI systems or nuclear magnetic resonance (NMR) systems, a static magnetic field B.sub.o is applied to the body under investigation to define an equilibrium axis of magnetic alignment in the region of the body under examination. An RF field is then applied in the region being examined in a direction orthogonal to the static B.sub.o field direction, to excite magnetic resonance in the region, and resulting RF signals are detected and processed. Generally, the resulting RF signals are detected by RF coil arrangements placed adjacent to the body. See for example, U.S. Pat. No. 4,411,270 to Damadian and U.S. Pat. No. 4,793,356 to Misic et al. Typically such RF coils are either surface type coils or volume type coils, depending on the particular application. Normally separate RF coils are used for excitation and detection, but the same coil or an array of coils may be used for both purposes. For multiple surface RF coils for use in NMR, see U.S. Pat. No. 4,825,162 to Roemer, et al., and for multiple volume RF coils for use in NMR, see U.S. Pat. No. 5,258,717 to Misic, et al.
A surface or volume RF coil is frequently used in examining the anatomy under investigation to obtain, for example, an image or spectroscopic or vascular data of the right symptomatic knee. The RF coil is placed very close to the symptomatic knee, adjacent the region to be imaged. Unfortunately, conventional RF coil arrangements suffer from the disadvantage that the RF coil picks up noise signals from around the region of interest (e.g., from the contralateral asymptomatic knee). This often results in a reduced overall signal-to-noise ratio (S/N) from the region of interest. Similar problems occur when imaging the breasts. Yet another example is when imaging the torso, with the arms close to the torso, or vice versa. Such problems are due in part to the fact that the range of the RF coil (i.e., the field of view (FOV) of the RF coil) covers a larger volume than the region of interest. In other words, the coil FOV is larger than the desired FOV. This is because the B field profile of the typical RF coil changes gradually in all directions due to the magnetic field properties.
Attempts have been made to alleviate such problems by way of flux rejection. See U.S. Pat. No. 5,382,903 to Young, and Burl and Young, Magnetic Resonance of Medicine, 36:326-330, 1996. The Young patent describes a mutually coupled resonant loop used to "buck" or reject the field generated by the receiving RF coil to a minimum outside the imaging FOV. Although the extent of the mutual coupling could be altered to a certain extent by changing the resonance of the flux rejecting loop, the induced currents depend primarily on the mutual coupling between coils which is a function of the physical separation between coils. The closer the respective coils the greater the coupling; and the greater induced currents resulted in improved RF screening. However, the opposite induced currents in the flux rejecting loop reduced the net signal intensity drastically over the imaging FOV of the coil. This extent varied with different coil spacings, which also affected the screening efficiency.
For example, FIG. 1a is a loop type coil 20 commonly used for several NMR applications. The coil 20 is resonated at the NMR frequency with three similar capacitors with values C1 and two capacitors with values 2C1. The rectangular loop serves as an inductance and resonates with a capacitance of approximately C1/4 at the NMR frequency. The center point between the two 2C1 capacitors is a virtual ground point VG and is forced to the real ground. This prevents any currents on the shields of the coaxial cables and therefore obviates the need for cable traps. The voltage across one of the 2C1 capacitors is matched to 50 ohms using a reactive network and fed to a preamplifier prior to system amplification and digitization (not shown).
The square box across the C1 capacitor represents a decoupling network 22 used for decoupling the RF coil during RF transmit. The number of decoupling networks 22 used in the RF coil 20 depends on the size of the RF coil and the degree of decoupling needed to minimize image artifacts and allow for safe preamplification operation. Additional circuitry may be used between the RF coil 20 and the preamplifier to reduce the amount of RF energy seen at the input of the preamplifiers during whole body RF transmit.
As shown in FIG. 11b, the decoupling network 22 consists of an inductor L1 and a pin diode D1. The pin diode D1 can be forward or reverse biased using programmable DC signals from the MRI system. Normally the pin diodes D1 will be forward biased such that the L1-C1 circuit is resonant at the NMR frequency. This will create a parallel trap, which will present ap open circuit with respect to the transmitting RF currents. During receive the pin diode D1 is reverse biased. Therefore, the L1-D1 circuit will effectively be open, and the RF coil will be resonant (with C1) at the NMR frequency.
FIG. 2 illustrates an RF loop coil 30 as described in the aforementioned Burl and Young article. Here, a secondary loop 34 is placed on one side of an RF coil primary 32. Both loops 32 and 34 are magnetically coupled through space, and hence are mutually coupled to one another. The primary loop 32 is resonant with capacitance C2 and the secondary loop 34 is resonant with a capacitor C3. Together the coil 30 is tuned to the NMR frequency. The capacitors in the primary and secondary loops are paralleled and include similar decoupling networks 22 as in FIG. 1. Although the mutual coupling between the primary and secondary loops can be varied to a certain extent by varying the resonance frequency of the secondary loop (by varying C3), the mutual coupling between the primary and the secondary loop 32 and 34 is primarily dominated by the magnetic coupling and hence their proximity to one another. The range of mutual coupling between the two loops by way of changing the resonance of the secondary loop is small. Although, a greater range may be obtained by physically moving the secondary loop, having moving coils in a RF coil package has not been a practical solution. Also, the closer the secondary loop the greater will be the induced currents. Since the currents induced would be opposite in direction from the RF coil primary, this will substantially reduce the net signal from the imaging FOV, which is not desirable.
Additional shortcomings associated with conventional RF coils have arisen due to limited control of the overlap between volume and surface type RF coils in the imaging FOV. Without the ability to control the extent of FOV overlap, particularly with regard to providing asymmetric overlap, a reduced usable imaging FOV results in a reduced filling factor which causes a reduced S/N. Furthermore, oftentimes it is desirable that an RF coil be tunable at two or more different NMR frequencies. However, conventional RF coils capable of being tuned at different frequencies either require traps for multiple tuning and/or result in significantly different FOVs at the different frequencies.
In view of the aforementioned shortcomings associated with conventional RF coil designs, there is a strong need in the art for an RF coil and method which provide the ability to vary the currents in the RF coil, such as to provide adequate RF screening, without significantly compromising S/N in the imaging FOV. Moreover, there is a strong need in the art for an RF coil arrangement and method which provides a high S/N over the imaging FOV by allowing an asymmetric overlap between volume and surface coils. In addition, there is a strong need in the art for an RF coil and method which provides for multiple tuning with similar B field profiles in the different NMR frequencies over the imaging FOV, without the use of traps.